The field of the invention is coherent imaging using vibratory energy, such as ultrasound, and, in particular, systems and methods for shear wave based elasticity imaging.
There are a number of modes in which ultrasound can be used to produce images of objects. For example, an ultrasound transmitter may be placed on one side of the object and sound transmitted through the object to an ultrasound receiver placed on the other side of the object. With transmission mode methods, an image may be produced in which the brightness of each pixel is a function of the amplitude of the ultrasound that reaches the receiver (“attenuation” mode), or the brightness of each pixel is a function of the time required for the sound to reach the receiver (“time-of-flight” or “speed of sound” mode). In the alternative, the receiver may be positioned on the same side of the object as the transmitter and an image may be produced in which the brightness of each pixel is a function of the amplitude or time-of-flight of the ultrasound reflected from the object back to the receiver (“reflection,” “backscatter,” or “echo” mode).
Acquisition of ultrasound data can be carried out with a number of backscatter methods. In the so-called “A-mode” method, an ultrasound pulse is directed into the object by an ultrasound transducer and the amplitude of the reflected sound is recorded over a period of time. The amplitude of the echo signal is proportional to the scattering strength of the reflectors in the object and the time delay is proportional to the range of the reflectors from the transducer. In the so-called “B-mode” method, the transducer transmits a series of ultrasonic pulses as it is scanned across the object along a single axis of motion. The resulting echo signals are recorded as with the A-mode method and their amplitude is used to modulate the brightness of pixels on a display. The location of the transducer and the time delay of the received echo signals locates the pixels to be illuminated. With the B-mode method, enough data are acquired from which a two-dimensional image of the reflectors can be reconstructed. Rather than physically moving the transducer over the subject to perform a scan it is more common to employ an array of transducer elements and electronically move an ultrasonic beam over a region in the subject.
The ultrasound transducer typically has a number of piezoelectric elements arranged in an array and driven with separate voltages. By controlling the time delay, or phase, and amplitude of the applied voltages, the ultrasonic waves produced by the piezoelectric elements (“transmission mode”) combine to produce a net ultrasonic wave focused at a selected point. By controlling the time delay and amplitude of the applied voltages, this focal point can be moved in a plane to scan the subject.
The same principles apply when the transducer is employed to receive the reflected sound (“receiver mode”). That is, the voltages produced at the transducer elements in the array are summed together such that the net signal is indicative of the sound reflected from a single focal point in the subject. As with the transmission mode, this focused reception of the ultrasonic energy is achieved by imparting separate time delays, or phase shifts, and gains to the echo signal received by each transducer array element.
Scanning of an object of interest using a transducer having an array of separately operable elements can be effectuated with electronic methods and systems such as, for example, linear array systems and phased array systems.
A linear array system includes a transducer having a large number of elements typically disposed in a line. A small group of elements are energized to produce an ultrasonic beam that travels away from the transducer, perpendicular to its surface. The group of energized elements is translated along the length of the transducer during the scan to produce a corresponding series of beams that produce echo signals from a two-dimensional region in the subject. To focus each beam that is produced, the pulsing of the inner elements in each energized group is delayed with respect to the pulsing of the outer elements. The time delays determine the depth of focus which can be changed during scanning. The same delay factors are applied when receiving the echo signals to provide dynamic focusing during the receive mode.
A phased array system commonly employs so-called phased array sector scanning (“PASS”). Such a scan is comprised of a series of measurements in which all of the elements of a transducer array are used to transmit a steered ultrasonic beam. The system then switches to receive mode after a short time interval, and the reflected ultrasonic wave is received by all of the transducer elements. Typically, the transmission and reception are steered in the same direction, θ, during each measurement to acquire data from a series of points along a scan line. The receiver is dynamically focused at a succession of ranges, R, along the scan line as the reflected ultrasonic waves are received. A series of measurements are made at successive steering angles, θ, to scan a pie-shaped sector of the subject. The time required to conduct the entire scan is a function of the time required to make each measurement and the number of measurements required to cover the entire region of interest at the desired resolution and signal-to-noise ratio. For example, a total of 128 scan lines may be acquired over a sector spanning 90 degrees, with each scan line being steered in increments of 0.70 degrees.
In the methods described above, multiple ultrasound beams are translated or steered to scan a two-dimensional (2D) area of an object of interest in order to form a 2D B-mode ultrasound image. Assuming that, for example, 100 ultrasound beams are used, the time required for forming such a 2D image is the aggregate of 100 transmit-receive events. If each transmit-receive event lasts 100 microsecond, then the formation of the 2D image takes about 10,000 microseconds. In another B-mode imaging method often referred to as “plane wave imaging”, a wave with a substantially planar wavefront is used in the transmission mode. This wave is not focused, typically has a spatial extent comparable to the total width of the aperture of the transducer, and can be formed by applying the transmit voltage to all elements of the transducer without time delay. Echo signals from this “plane wave” transmission are received by each transducer element and stored in corresponding tangible data storage. The stored echo signals from multiple transducer elements are delayed and summed together to reconstruct the ultrasound wave reflected from the object at the location of a given single pixel in the 2D image. This process is repeated for each pixel to obtain the “focused” echoes from all pixels to form an overall 2D image. Different delay parameters are introduced to focus at different pixels. In “plane wave imaging”, only one transmit-receive event is required to form a 2D image. Therefore, the image acquisition time for one frame (one 2D image) is small, leading to a higher frame rate as compared to the beam scanning methods described above. Assuming that a transmit-receive event lasts 100 microsecond, for example, a plane wave imaging procedure can have a frame rate of about 10 kHz, whereas the aforementioned example of 100-beam scanning method has a frame rate of 100 Hz.
The same scanning methods may be used to acquire a three-dimensional image of the subject. The transducer in such case is a two-dimensional array of elements which steer a beam throughout a volume of interest or linearly scan a plurality of adjacent two-dimensional slices.
Characterization of mechanical properties of the tissue, particularly the elasticity or tactile hardness of tissue, has important medical applications because these mechanical properties are closely linked to tissue state with respect to pathology. For example, breast cancers are often first detected by the palpation of lesions with abnormal hardness. In another example, a measurement of liver stiffness can be used as a non-invasive alternative for liver fibrosis staging.
The radiation force of ultrasound can be used to generate, remotely, a shear wave within the tissue for noninvasive elasticity imaging. Traditionally, a focused ultrasound beam with long duration (for example, a few hundred microseconds, containing many ultrasound cycles) is used to impart tissue motion at the focus of the ultrasound push beam (referred to as push origin), and a pulse echo ultrasound is used to detect the shear wave propagating outwards from the push origin. The shear wave propagation speed is measured and used to estimate viscoelastic properties of tissue.
The motion of the tissue caused by the push beam is typically very weak (on the order of micrometers), which undermines the reliability of the detection of the shear wave and estimation of tissue viscoelastic properties. The tissue motion increases with amplitude of the push ultrasound beam. However, FDA requirements limit the Mechanical Index (MI) of diagnostic ultrasound to below 1.9 for diagnostic applications in human species. Therefore, the amplitude of the used push beam cannot exceed a certain threshold to avoid exceeding the MI limit.
Another way to increase amplitude of the tissue motion includes increasing the duration of a push pulse. Due to limitations of hardware (for example, due to power droop of transmission circuits) or software (for example, ultrasound output safety watchdogs), current commercial ultrasound scanners may not be equipped to form a push beam with long enough duration to produce a shear wave having sufficiently high amplitude. It would therefore be desirable to provide a method for generating large amplitude shear waves in tissues and do so in a manner that remains within FDA safety limits and that is within the capacity of commercial ultrasound scanners.
Current methods for shear wave detection and shear wave speed estimation are designed to suit traditional ultrasound scanners which image one line from one transmit-receive event. Typically, a shear wave is detected at only a few positions defined along the shear wave propagation path, and detection at each of these positions is repeated at a Pulse Repetition Frequency (PRF) of several kilohertz for tens of milliseconds. Such relatively long detection period may limit the practically use of currently employed methods in responding or detecting rapid changes of tissue elasticity (in a beating heart, for example). In addition, detection of a shear wave signal during the long detection period inevitably exposes the detection process to interference from physiological tissue motion (for example, gross motion of the beating heart). It would therefore be desirable to provide a method for estimation of shear wave speed that requires detection of shear wave in a shorter time period.